In an MRI apparatus, a magnet has been conventionally employed for generating a static magnetic field, the magnet comprising a narrow tubular solenoid coil, which can efficiently generate a uniform static magnetic field. However, in the MRI apparatus of this type, the object to be examined is placed in a space like a tunnel, imparting an oppressive feeling to the object. It is also known that such a condition causes fear in a claustrophobic person and children. Accordingly, an open MRI apparatus employing a flat type gradient magnetic field coil, in which one side of the magnetic device is closed off and the front of the magnet is opened up to widen the space for the object to enter, has recently seen increasing use.
Since such an open MRI apparatus, using an open magnet, has an air space to enable an interventional procedure in which a doctor performs treatment while observing an image, advanced medical institutions have been employing this type of apparatus in recent years. In such an MR-interventional procedure, it is a requirement that the operator be able to check the MR image of the procedure being carried out in real time. To improve the image quality and the functions of real-time imaging (high-speed imaging), the MRI apparatus requires a gradient magnetic field coil that operates at high speed and a suitable driving power for this gradient magnetic field coil, as well as an RF coil for detecting NMR signals with high sensitivity, and a magnet having a high static magnetic field strength.
A magnet having a high static magnetic field strength as required for an open MRI apparatus is proposed in Japanese Patent Laid-open Publication JP-A-10-179546, JP-A-11-155831, and JP-A-11-197132.
As for the gradient magnetic field coil, on the other hand, a switching power having large capacity has been developed, and thus the gradient magnetic field coil can be driven at high-speed. However, an electromagnetic force works on the gradient magnetic field coil when current for driving the gradient magnetic field is applied in pulse form, and so mechanical distortion and vibration tend to occur, and this vibration is transmitted to the magnet for generating the static magnetic field. Thus, some problems occur in that the uniformity of the magnetic field in the measurement space is deteriorated and the acoustic noise due to the vibration of the gradient magnetic field coil is increased.
The above-mentioned problems are far more serious in an MRI apparatus having an open-gantry structure. That is, the open MRI apparatus is more liable to be affected by vibration than an apparatus employing a tunnel type magnet. The vibration transmitted from the gradient magnetic field coil to the magnet deteriorates the stability of the static magnetic field generated by said magnet, and thus the image quality is deteriorated.
Further, as described above, both noise and vibration increase during high-speed imaging in real time, since the gradient magnetic field coil is subject to a rapid switching-drive. For example, when performing an imaging procedure, such as an EPI method in which the gradient magnetic field coil is subject to rapid switching-drive, the operating noise of the gradient magnetic field can be over 100 dB.
Conventionally, various techniques as described below have been proposed for solving the above-mentioned problems concerning the noise and vibration of the gradient magnetic field coil: it has been proposed (a) to increase the strength of the gradient magnetic field coil bobbin; (b) to limit the amplitude of the vibration by increasing the weight of the gradient magnetic field coil; (c) to insert a number of lead balls in the structure of the gradient magnetic field coil to absorb the vibration energy of the coil by converting it into frictional heat generated between said lead balls; and (d) to reduce the noise of the gradient magnetic field by generating sound having a reverse phase to that of the noise.
Each of the above-described methods has achieved about a 10 dB noise reduction. However, in spite of the amount of noise reduction, those methods have created some other problems, that is, the weight of the gradient magnetic field coil is increased and the control procedure becomes complicated when using those methods.
Concerning such problems, an MRI apparatus employing a magnet of the solenoidal coil type, in which the noise reduction effect is improved, is proposed in Japanese Patent Laid-open Publication No. JP-A-10-118043. In this MRI apparatus, the gradient magnetic field coil is covered with equipment that can vacuumize said coil, and it is not placed on the magnet but is independently installed on the magnet-installing floor. According to this method, both the air-propagated vibration resulting from the driving of the gradient magnetic field coil and the vibration that is propagated through fixed objects into the magnet can be reduced. Consequently, the noise of the whole apparatus resulting from the vibration of the gradient magnetic field coil can be reduced by 20-30 dB.
However, it is difficult to apply this method to an open MRI apparatus; that is, in the open MRI apparatus, the gradient magnetic field coil is divided into an upper part and a lower part, and they are arranged near the facing plane of the magnet. Therefore, a method of fixing the upper part of the coil that can secure the required openness has to be devised.
An MRI apparatus has problems of reducing vibration and noise of the gradient magnetic field coil, as well as of reducing eddy currents generated on the surface of the magnet vessel placed near the gradient magnetic field coil when the pulse current is transmitted to said coil.
In an MRI apparatus, the gradient magnetic field coil is placed very close to the magnet vessel so as to secure as wide an examination space as possible. Thus, when a pulse current is applied to said gradient magnetic field coil, a sudden flux change occurs in the space surrounding said magnet vessel with each rise and fall of the pulse. This flux change causes eddy currents to be generated on the surface of the magnet vessel placed near the gradient magnetic field coil. The direction of said eddy currents generated on the surface of the magnet vessel is opposite to that of the gradient magnetic field generated by said gradient magnetic field coil, and thus they hinder the rising and falling of the gradient magnetic field. This effect causes a variance in the amount of the gradient magnetic field which must be applied and interferes with execution of the high-speed imaging.
Therefore, methods to prevent such an effect have been conventionally applied to the apparatus. One of them is to put an active shield coil on the gradient magnetic field coil. When the active shield type gradient magnetic field coil is used in the open MRI apparatus, a pair of the gradient magnetic field coils is arranged above and below the examination space, and behind them a pair of shield coils is placed. Here, the gradient magnetic field coils have axes respectively in the x-, y-, and z-axis directions so as to generate gradient magnetic fields in these three directions in the examination space.
Correspondingly, the shield coil also has axes respectively in the x-, y-, and z-axis directions.
The relation between the gradient magnetic field coil and the shield coil will be described with reference to the drawings to explain it simply. FIG. 28 shows a gradient magnetic field coil 610b arranged in the x direction, a shield coil 620b arranged in the x direction, and the magnet vessel 21b. Current in the direction of the arrow M1 (counterclockwise) and in the direction of the arrow M2 (clockwise) is applied to the gradient magnetic field coil 610b, and current in the direction of the arrow S1 (clockwise) and in the direction of the arrow S2 (counterclockwise) is applied to the shield coil 620b. At this point, magnetic fields in the direction of the arrow U1 (upward) and in the direction of the arrow D1 (downward) are generated by the gradient magnetic field coil 610b, and magnetic fields in the direction of the arrow D2 (downward) and in the direction U2 (upward) are generated by the shield coil 620b. When those gradient magnetic fields are combined with the gradient magnetic field generated by the gradient magnetic field coil symmetrically placed across the examination space, an upward gradient magnetic field, indicated by the arrow UX, is generated in the positive region in the x direction, and a downward gradient magnetic field, indicated by the arrow DX, is generated in the negative region in the x direction.
FIG. 29 is a graph showing the calculated magnetic field leakages of the coils 610b and 620b on the observation line 650, shown by a broken line on the magnet vessel 21b. Referring to FIG. 29, the horizontal axis represents the distance from the center of the magnet vessel, and the vertical axis represents the strength of the magnetic field leakage. The curved line M indicates the magnetic field leakage due to the coil 610b, the curved line S represents the magnetic field leakage due to the coil 620b, and the curved line T represents the composite value of those magnetic field leakages. In the more detailed illustration of said curved line T, as shown in FIG. 30, the magnetic field leakage is canceled by both of those coils in the region near the magnet vessel. However, it suddenly becomes large at a certain distance away from the center, that is, at the position on the outside of the radius of the gradient magnetic field coil. The reason for this may be that the coil conductor of the shield coil is arranged over a wider range than the main coil conductor of the gradient magnetic field in order to increase the shielding effect, and the magnetic field strength on the surface of the magnet vessel is inversely proportional to the square of the distance from the surface of the cryostat of the coil generating the magnetic field.
FIG. 31 shows how the gradient magnetic field coil 610a and the shield coil 620a are arranged relative to the surface of the cryostat and the z direction, using a cross-sectional view along the observation line 650. The shield coil 620a and the gradient magnetic field coil 610a are arranged below the cryostat 21a. For example, six conductors are arranged on the shield coil surface 410 of the shield coil 620a, and eight conductors are arranged on the main coil surface 400 of the gradient magnetic field coil 610a. The coil conductors of the shield coil 620a are spaced more extensively than the coil conductors of the gradient magnetic field coil 610a. When the pulse current is applied to these coils, a magnetic field distribution is generated by the coil conductors near the outer circumference of the coils on the surface of the cryostat 21a, as illustrated in FIGS. 32(a) to 32(e).
Referring to FIG. 32(a), 610a designates the coil conductors of the gradient magnetic field coil, 630 designates the distribution of the magnetic field generated by one coil conductor of the coil 610a, and 640 designates the composition of the magnetic field distributions generated by a plurality of the coils 610a. Referring to FIG. 32(b), 620a designates the coil conductors of the shield coil, 650 designates the magnetic field distribution generated by one of the conductors of the coil 620a, and 660 designates the composition of the magnetic field distributions generated by a plurality of conductors of the coil 620a. FIG. 32(c) shows the difference between these magnetic distributions 640 and 660. The parts shaded with oblique lines are the magnetic fields which are not canceled, and this magnetic flux is leaked to the cryostat 21a. Consequently, the magnetic field indicated with the oblique lines leaks to the magnet vessel, and thus an eddy current is generated on the surface of the cryostat 21a. 
The present invention is directed to the aspects and problems described above. The first object of the present invention is to secure sufficient air space around the magnet, and to reduce vibration of the gradient magnetic field coil resulting from the driving of gradient magnetic field and the noise accompanying said vibration in an open MRI apparatus.
The second object of the present invention is to provide an open MRI apparatus in which the uniformity of the static magnetic field is not disturbed due to the vibration of the gradient magnetic field coil transmitted to the static magnetic field generating magnet.
The third object of the present invention is to provide an open MRI apparatus in which comfortable conditions for examination can be provided to the object, and in which the MR-interventional procedure using high-speed imaging can be performed with low noise.
Further, the fourth object of the present invention is to provide an active-shielded gradient magnetic field coil for reducing magnetic field leakage generated at the edge area of the gradient magnetic field coil when a pulse current is applied to the gradient magnetic field coil, that is, for reducing eddy currents generated on the surface of the magnet vessel, as well to provide an open MRI apparatus using said active-shielded gradient magnetic field coil by which an excellent image can be obtained.